Gamma ray cameras are well-known devices used to image the distribution and concentration of a radioactive field by detecting gamma ray emissions from radioactive decay. The conventional gamma or “Anger” camera (named after its inventor) is described in U.S. Pat. No. 3,011,057 for RADIATION IMAGE DEVICE, hereby incorporated by reference. The gamma camera typically uses a sodium iodide (“NaI”) scintillation crystal to detect gamma ray emissions from a radioactive object. The scintillation crystal is positioned to receive a portion of the gamma ray emissions from the decay of a radioactive isotope within the object. When a gamma photon strikes and is absorbed in the detector crystal, the energy of the gamma ray is converted into a large number of scintillation light photons that emanate from the point of the gamma ray's absorption in the crystal. This basic structure can be used for many different imaging studies, including PET, SPECT and planar imaging.
The gamma camera typically uses an array of photomultiplier tubes (PMTs), optically coupled to the crystal, which detect a fraction of these scintillation light photons and produce an electronic signal that is proportional to the number of incident scintillation light photons detected. The signals from the different photomultiplier tubes in the array are combined to provide an indication of the position and energy of gamma rays incident on the crystal.
Alternatives to the conventional photomultiplier tubes have long been sought for nuclear medical applications, in order to improve image quality and performance characteristics. In particular, photomultiplier tubes are relatively large, bulky, sensitive to magnetic fields, susceptible to linearity distortions, suffer from relatively low quantum efficiency, require a high supply voltage, and are subject to “dead” space between tubes in an array due to the inherent restrictions in the geometric shapes of the photomultiplier tubes and consequent limits in packing density and spatial resolution.
In particular, gamma cameras with solid-state detectors are known in the art. See, e.q., U.S. Pat. Nos. 4,055,765, 6,242,745, 6,359,281 and 6,921,904. Such solid state detectors take the place of the scintillation crystal and PMT, as the gamma photons are directly absorbed in the semiconductor material and the resultant induced electrical charges are measured at output terminals of the semiconductor detectors. However, such solid state detectors require expensive cooling systems because of the significant heat generated by the absorption of gamma photons in the semiconductor material and resultant electric charge produced therein.
Also known are gamma cameras having photodiode detectors instead of PMTs, coupled to a scintillation crystal. See, e.q., U.S. Pat. Nos. 4,234,792, 5,171,998, and 5,773,829. However, efforts to commercialize such photodiode detectors generally have not been successful, as a result of performance-related issues such as insufficient amplification, lack of stability, and large capacitance.
Therefore, there remains a need in the art for improved an photodetector for use in radiation imaging apparatus such as PET, SPECT, and CT medical applications, as well as other non-medical radiation detection applications.